One common type of detector for nuclear medicine imaging systems is the scintillation detector. In this detector gamma-rays or other ionizing radiations are absorbed by high density scintillator materials such as lutetium oxyorthosilicate (LSO), lutetium-yttrium oxyorthosilicate (LYSO), sodium iodide (NaI) or cesium iodide (CsI), for example. The energy deposited in the scintillator by the gamma-rays is converted into visible light photons, which are subsequently detected by a photodetector coupled to the scintillator. Photodetectors commonly used in nuclear medicine include photomultiplier tubes (PMTs), avalanche photodiodes (APDs) and Geiger-mode APDs (GAPDs), also called silicon photomultipliers (SiPMs) or multi-pixel photon counters (MPPCs).
Devices for detecting the distribution of gamma rays transmitted through or emitted from objects to study the compositions or functions of the objects are well known to the art, e.g. the techniques referred to as Emission Computed Tomography can be divided into two specific classes; Single Photon Emission Computed Tomography (SPECT) uses radiotracers which emit gamma-rays and Positron Emission Tomography (PET) which uses radiotracers that emit positrons. Therefore, the fundamental physical difference between the two techniques is that PET uses annihilation coincidence detection to define a line of response (LOR) while SPECT uses an absorptive collimator to define the directionality of the detected gamma-ray. The PET technique can determine, in-vivo, biochemical functions, on the injection of biochemical analog radiotracer molecules that emit positrons in a living body. The positrons annihilate with surrounding electrons in the subject body to produce a pair of annihilation photons (also sometimes referred to as gamma-rays), each having 511 keV of energy; traveling in nearly opposite directions. The detection of a pair of annihilation photons or gamma-rays by two opposed detectors allows for the determination of the location and direction in space of a trajectory line defined by the opposite trajectories of the gamma-rays. Tomographic reconstruction is then used to superpose the numerous trajectory lines obtained by surveying the subject with an array of detectors to image the distribution of radiotracer molecules in the living body.
Spatial resolution of a nuclear medicine imaging detector depends on how precisely the location of the gamma-ray interaction in the scintillator crystal can be determined. One commonly used approach to create high resolution detectors is to pixilate the scintillator crystal into small elements. In this type of design, the spatial sampling of the system is determined by the size of the scintillator crystal element used. These small crystal sizes make it difficult to use a one-to-one coupling (scintillator crystal to photodetector) due to the small size of photodetector required and the large number of readout channels required.
An alternative to one-to-one coupling is to use a light-sharing design in which an array of scintillator crystals are read out by a small number of photodetectors. In these detectors, the location of the gamma interaction (i.e. determination of which scintillator crystal the gamma interacted in) is measured by comparing the relative signals in the array of photodetectors. These light-sharing designs are in general based on the original ‘block detector’ design.
Light sharing detectors typically require a light guide to be placed between the scintillator crystal array and the photodetector array in order to sufficiently distribute the light between the sensors in the photodetector array and allow an accurate determination of position of interaction. If there is not a sufficient amount of light-spread, then two neighboring scintillator crystals might produce signal patterns on the photodetectors that are too similar to be individually resolved. Without the light guide, there would not be sufficient light spread to allow the decoding of these small crystals, particularly for crystals along the perimeter of the scintillator crystal array and the photodetector array.
These light guides can be made of a glass or other optically transparent material with a thickness of up to a few millimeters for silicon photomultiplier (SiPM) based photodetector arrays. In many cases, the active area of the photodetector array is less than the total package area of the photodetector array, creating a dead-space around the photodetector array. For scintillation detector manufacturing it is desirable to have a scintillator crystal array that is the same area as the total photodetector package area to avoid gaps between adjacent detector modules created by the dead space around the photodetector arrays. In these cases, it is common to use a tapered or similar design light guide that minifies the footprint of the scintillator crystal array so that it can be read out by the photodetector array with smaller area. This technique is used successfully in many photomultiplier tube (PMT) based detector designs.
SiPM based photodetector arrays are being introduced that area created by tiling individual photodetector pixels, each with very minimal dead-space borders, leading to photodetector arrays that have nearly zero dead-space around their border. The lack of dead-space means that the scintillator crystal array can have an active area equal to the total photodetector array area and there is no need for a minifying light guide. Since the photodetector array is assembled from individual photodetector pixels that are not optically contiguous, there is still a need for a light guide between the scintillator and photodetector arrays to act as a light diffuser for the purpose of spreading light between multiple photodetector array elements.
In detectors for nuclear imaging, particularly in applications involving positron emission tomography (PET), it is desirable to have detectors with the capability to determine the depth of the gamma-ray interaction along the length of the scintillator element. Detectors with this ability are termed depth-of-interaction (DOI) capable detectors.